Method for filtering reflexes in full-field setups for ophthalmologic imaging by separated illumination and detection apertures

ABSTRACT

A parallel detecting optical coherence tomography (OCT) setup and method, in which the light paths of the illumination of the sample and of the detection of the backscattered light do not use the same apertures. The separation of illumination and detection apertures filters these disturbing reflexes from the backscattered light of the sample and significantly increases image quality.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.14/610,319 filed on Jan. 30, 2015, which claims the benefit of U.S.Provisional Patent Application No. 61/934,265 filed on Jan. 31, 2014.The disclosures of U.S. patent application Ser. No. 14/610,319 and U.S.Provisional Patent Application No. 61/934,265 are hereby incorporated byreference.

FIELD OF THE INVENTION

The present invention relates to the field of ophthalmologic imaging,and more particularly to systems and methods for filtering reflexes infull-field setups for ophthalmologic imaging by separated illuminationand detection apertures.

BACKGROUND

In ophthalmology tomographic imaging of scattering structures of the eyeis of high interest. The standard technique is optical coherencetomography (OCT), employing an interferometric setup and a spatiallycoherent light source with a short temporal coherence length.Conventional OCT systems acquire volumetric data by scanning a focusedbeam over the sample and consequently, measurement speed is limited bythe scanning speed. To further increase imaging speed methods withparallel detection of scattered light have been developed. Since inthose tomographic ophthalmic imaging techniques that employ paralleldetection no confocal imaging is used, disturbing reflexes reduce theimage quality significantly.

Below are some existing ophthalmology tomographic imaging techniques:

In Fourier-Domain optical coherence tomography (FD-OCT) an interferencesignal over a broad spectral width is recorded in an interferometricsetup. This is achieved by either detecting the signal of a broadbandlight source spectrally resolved (Spectral-Domain OCT) or by recordingan interference signal over time, while a laser source is spectrallytuned (Swept-Source OCT). Considering the entire spectral range, theshort temporal coherence length of the light source allows for anoptical path length measurement of backscattered and/or reflected lightfrom a sample in one of the interferometer arms. The path length isencoded in the interference signal, which is generated by superimposinglight from sample and reference arm. Measurements of interferencesignals at several wavelengths allow a depth encoded profile of thesample (A-scan). The main area of application for OCT is ophthalmologicimaging, especially tomographic images of the retina (posterior eyesegment) and structures in the anterior eye segment (cornea, ocularlens, iridocorneal angle).

Swept-Source OCT (SS-OCT) uses a setup with a tunable light source.While tuning the light source a wavelength-dependent interferencespectrum is measured. The tuning range of the laser, which correspondsto its entire spectrum defines the coherence length of the system. Thecoherence length of a single wavelength is defined by the instantaneousline width. A tomographic data volume is measured conventionally bysequential data acquisition at different points, i.e., lateral scanningof the sample. Generally the sample is screened by two lateral scannersand the backscattered light is detected with a point detector.Measurement speed for sequential data acquisition might be limited bythe speed of the scanners in this scenario. A phase stable detectionover the whole volume is not possible in most cases. Scanning fiberbased OCT systems have a confocal gating, i.e. only sample structures inthe focus are illuminated and detected, while out-of-focus photons arerejected. Therefore multiple scattered photons occurring in stronglyscattering media are suppressed.

A setup with a partly parallelized data acquisition is calledLine-Field-OCT or Swept-Source parallel OCT (see reference 1). Thebackscattered light from the sample is detected in parallel in onelateral dimension, while it is scanned in the other lateral direction.The detector consists of multiple individual elements, which arearranged in a line, e.g., a line scan camera. Measurement speed isgenerally improved by the partial parallelization. Multiple scatteredphotons are not suppressed, if their last scattering event is recordedby the detector.

For a completely parallel detection in two dimensions the sample isilluminated homogeneously and spatially coherent (Full-FieldSwept-Source OCT, short: FF-SS-OCT). An area scan camera is used asdetector. In FF-SS-OCT the sample or rather a part of the sample isimaged onto the area camera. An advantage of this method is theincreased measurement speed. While in scanning OCT systems the scannersor the tuning speed limit the measurement speed, in FF-SS-OCT the framerate of the camera is in general the most limiting factor. All paralleldetected A-scans are phase stable to each other. With this methodscattered photons from all depths are detected, which are filtered inscanning systems by the confocal gating. Thus the measurement depth islarger in comparison to scanning systems, but the lateral resolutiondegrades out-of-focus. This degradation limits the useful measurementdepth, especially at high lateral resolution. Another disadvantage ofFF-SS-OCT is the detection of multiple scattered photons due to theparallel detection. Ophthalmic imaging with FF-SS-OCT has beendemonstrated successfully, showing in vivo retina measurements (seereference 2)

Holoscopy is method related to FF-SS-OCT where the sample is notnecessarily imaged onto the camera, but wave fields of the backscatteredlight from the sample are detected (see references 3 and 4). Thefocusing in all depth is performed in the following reconstructions.This has the advantage, that the lateral resolution does not degradeout-of-focus, but is constant over the whole volume. The reconstructionalgorithm for holoscopy is also suitable for increasing the focus depthin FF-SS-OCT data.

By implementing an off-axis reference illumination in FF-SS-OCT orholoscopy it is possible to separate the signal of the interference ofthe sample with the reference light from its complex-conjugated signalas well as from DC and autocorrelation signals and to suppress thenon-relevant signal terms. This increases the sensitivity of the imagingand avoids an overlay of multiple signal parts. (See reference 5).

All OCT techniques and related imaging methods mentioned so far are inparticular suitable for scattering samples. If a sample has highlyreflecting and weakly scattering parts, the strong reflections induceoverexposure artifacts in the images and decrease the sensitivity of themeasurements. More significantly, strong reflexes that are not withinthe measurement range, induce an incoherent background noise on alldepth profiles/A-scans.

In other ophthalmologic imaging modalities, which have spatialincoherent illumination, strong reflexes induce image artifacts as well.This is the case for fundus cameras, where photographs of the posterioreye segment are taken, as well as for silt lamps, where all segments ofthe eye can be visualized enabling a variable slit shaped illumination.The artifacts caused by strong reflexes mainly lead to an overexposureof the detector, or the reflexes overshadow the actual image of theretina. The approaches to reduce those artifacts are based on theseparation of illumination and detection apertures.

In conventional fundus camera setups a ring shaped aperture in theillumination light path is imaged into the plane of the pupil(References 6, 7, 8). The ring shaped illumination generates onlyreflexes in the outer areas, which are reflected at an angle, in a way,that they are not imaged onto the detector. The retina is illuminateddivergently. The backscattered light is refracted via the opticalelements of the eye and projected onto the detector. Such an aperture isnot possible with coherent light, as the light will interfere on theretina and not create a constant illumination.

In conventional slit lamp setups the eye is illuminated at an angle. Theangle is adjusted in a way that the backscattered light is detected,while the reflected light does not reach the detection light path(references 9, 10). When using a coherent light source the illuminationof the retina is not uniform due to interference effects.

Both fundus cameras and slit lamps use conventional light sources, e.g.,filament lamps or halogen bulbs. Those light sources are both spatiallyand temporally incoherent. Therefore there are no additionalinterference effects, which prevent a uniform illumination of thesample. In addition, these techniques do not provide depth informationfor scattering tissue.

So far there is no imaging modality that provides volumetric sampleinformation with a parallel detection and spatially coherent light thatimplements reflex reduction methods to increase SNR by decreasing straylight.

As described above, existing techniques have drawbacks anddisadvantages. Therefore, there is a need for an imaging system andmethod that allows imaging a larger portion of the eye than possiblewith any of today's optical imaging systems. Anterior and posteriorsegment can be imaged in parallel. Using parallel detection, volumetricimaging equivalent to >5 million A-scans per second (comparing toscanned OCT technology) are possible and provide very fast acquisition.The parallel detection scheme has a much higher efficiency, as it missesthe confocal gating that is used in the prior art of eye inspectionusing optical coherence tomography. Hence a larger imaging speed can beapplied within the restrictions of eye safety regulations.

SUMMARY

The present invention provides a parallel detecting OCT setup and methodwhere the light paths of the illumination of the sample and of thedetection of the backscattered light do not use the same apertures. Theseparation of illumination and detection apertures filters thesedisturbing reflexes from the backscattered light of the sample andsignificantly increases image quality. Similar approaches have beenimplemented in non-tomographic ophthalmologic imaging modalities withspatially incoherent light sources (mainly fundus cameras, slit lamps).However, different technical constraints apply and special technicalarrangement have to be used for tomographic imaging. For example, thelight sources that are used in fundus imaging are white light sources(such as a halogen lamp) which are incoherent. For tomographic imaging,spatially (and temporally) coherent radiation (a swept source laser) isrequired to obtain the depth information. Using the previous approaches(such as ring illumination) to separate the illumination from thedetection aperture results in disturbing interference patterns on theretina if coherent radiation is used, whereas everything is fine withincoherent light. This poses a problem when using the ring aperture fortomographic imaging, as the areas with destructive interference cannotbe imaged. Therefore, the approaches provided by the present inventionhave a unique advantage that it does not cause interference signals onthe retina.

One embodiment of the present invention provides an ophthalmologicimaging system including: a tunable light source; a first optical systemto focus light from the light source onto a mirror, such that the mirrorreflects the light onto a sample in a first light path; a second opticalsystem configured to direct scattered light from the sample onto adetector in a second light path, and the second optical system isfurther configured to direct a reference beam onto the detector, thereference beam being either on-axis or off-axis to the second path;wherein the mirror is configured to reflect the light in the first lightpath, which is an angle to the second light path.

Another embodiment of the present invention provides an ophthalmologicimaging system including: a tunable light source; a mirror with a hole;a first optical system to focus light from the light source into thehole of the mirror, such that the light passes through the hole isfocused onto the cornea by a lens and illuminates the retinadivergently; a second optical system configured to direct scatteredlight from the sample that is reflected by the mirror onto a detector,and the second optical system is further configured to direct areference beam onto the detector, the reference beam being eitheron-axis or off-axis to the second path.

Another embodiment of the present invention provides an ophthalmologicimaging system including: a tunable light source; a mirror with a hole;a gradient-index (GRIN) lens placed in the hole; a first optical systemto direct light from the light source onto the GRIN lens, such that thelight passes through the GRIN lens is focused onto the cornea by a lensand illuminates the retina divergently; a second optical systemconfigured to direct scattered light from the sample that is reflectedby the mirror onto a detector, and the second optical system is furtherconfigured to direct a reference beam onto the detector, the referencebeam being either on-axis or off-axis to the second path.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a ophthalmologic imaging setup in accordance with anembodiment of the invention.

FIG. 2 is a ophthalmologic imaging setup in accordance with anotherembodiment of the invention.

FIG. 3 a ophthalmologic imaging setup in accordance with yet anotherembodiment of the invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The description of illustrative embodiments according to principles ofthe present invention is intended to be read in connection with theaccompanying drawings, which are to be considered part of the entirewritten description. In the description of embodiments of the inventiondisclosed herein, any reference to direction or orientation is merelyintended for convenience of description and is not intended in any wayto limit the scope of the present invention. Relative terms such as“lower,” “upper,” “horizontal,” “vertical,” “above,” “below,” “up,”“down,” “top” and “bottom” as well as derivative thereof (e.g.,“horizontally,” “downwardly,” “upwardly,” etc.) should be construed torefer to the orientation as then described or as shown in the drawingunder discussion. These relative terms are for convenience ofdescription only and do not require that the apparatus be constructed oroperated in a particular orientation unless explicitly indicated assuch. Terms such as “attached,” “affixed,” “connected,” “coupled,”“interconnected,” and similar refer to a relationship wherein structuresare secured or attached to one another either directly or indirectlythrough intervening structures, as well as both movable or rigidattachments or relationships, unless expressly described otherwise.Moreover, the features and benefits of the invention are illustrated byreference to the exemplified embodiments. Accordingly, the inventionexpressly should not be limited to such exemplary embodimentsillustrating some possible non-limiting combination of features that mayexist alone or in other combinations of features; the scope of theinvention being defined by the claims appended hereto.

This disclosure describes the best mode or modes of practicing theinvention as presently contemplated. This description is not intended tobe understood in a limiting sense, but provides an example of theinvention presented solely for illustrative purposes by reference to theaccompanying drawings to advise one of ordinary skill in the art of theadvantages and construction of the invention. In the various views ofthe drawings, like reference characters designate like or similar parts.

Ophthalmologic imaging is sensitive to image artifacts that are causedby reflections from the eye itself. Especially the front side of thecornea is strongly reflecting. In swept-source OCT and relatedtomographic imaging methods these reflections cause a reducedsensitivity due to overexposure, coherent background, and/or incoherentbackground noise. By implementing separated apertures for theillumination of the eye and the detection of the backscattered lightstrong reflections at interfaces like the cornea can be filtered beforedetection. In one embodiment of the invention, an interferometric setupfor ophthalmic imaging includes one or more of the following components:

-   -   A tunable light source; preferably, the total coherence length,        defined by the spectral width of the tuning range, does not        exceed 50 μm. The instantaneous coherence length, defined by the        instantaneous line width of the spectral range during one shot,        preferably, is at least 200 μm.    -   A detector consisting of several pixels, preferably, recording        data with an acquisition speed of at least 500 images per        second.    -   A reference illumination with an incident direction that can        differ from the incident direction of the sample light on the        detector (off axis).    -   A spatial coherent or partly coherent illumination of the eye.    -   A separated illumination and detection aperture for the        suppression of reflected light from plane or curved interface,        preferably, of at least 10 dB.

Separated illumination and detection apertures in an interferometricsetup could be implemented as follows:

In one embodiment of the invention, the angle under which the sample isilluminated differs from the detection angle of the backscattered light,as illustrated by FIG. 1. A light source 110 is used to illuminate thesample 130. The illumination light from light source 110 is focused by acollimator 120 and a lens 180 onto a mirror 140. The mirror reflects theillumination light into the sample 130. The backscattered light from thesample 130 is detected by a camera 170. A reference signal 150 (whichcan be either on-axis or off-axis) is directed onto the camera 170 via abeam splitter 160. As shown in FIG. 1, the illumination light path isreflected by the mirror 140, which is placed at an angle to thedetection light path. The angle of the mirror is adjusted in a way thatthe desired area of the sample 130 is illuminated and the backscatteredlight detected while the reflexes of the cornea are not in the detectionbeam path.

In another embodiment of the invention, central reflexes of the corneacan be filtered by creating a focus on the cornea or slightly beneaththe cornea surface. The illumination of the retina is divergent.Backscattered light from the retina is detected while reflected lightfrom the cornea is filtered, as shown in FIG. 2. A light source 210 isused to illuminate the eye 230. A collimator 220 and a lens (280) focusthe illumination light path through a central hole in an angled mirror240. Further optical elements create a focus on the cornea, while theretina is divergently illuminated. The backscattered light is convertedinto the far field by the optics of the eye. The central cornea reflexdoes not reach the detection camera, since it is imaged back through thehole in the mirror and/or is absorbed by a non-reflecting area aroundthe hole, while most of the light from the retina is deflected by themirror 240 onto the detector 260. Only a small part of the sample lightis filtered by the mirror. A reference signal 250 (which can be eitheron-axis or off-axis) is directed onto the camera 260 via a beam splitter270.

Another option to create a focus on the cornea is to position aGradient-index (GRIN) lens with a suitable pitch length in a hole in anangled mirror, in accordance with one embodiment of the invention. FIG.3 shows a GRIN lens 380 in a hole of the mirror 340. The illuminationbeam is converging when leaving the GRIN lens 380 and focused by anadditional lens onto the cornea or slightly beneath the cornea surfaceof the eye 330. The backscattered light is converted into the far fieldby the optics of the eye. The central cornea reflex does not reach thedetection camera, since it is imaged back through the hole in the mirrorand/or is absorbed by a non-reflecting area around the hole, while mostof the light from the retina is deflected by the mirror 340 onto thedetector 360. Only a small part of the sample light is filtered by themirror. An off-axis reference signal 350 is directed onto the camera 360via a beam splitter 370. FIG. 3 also shows an arrangement according toan embodiment in which a light source 310 supplies both the illuminationbeam and reference beam. A coupler 390 that splits the light from lightsource 310 into two paths, one to the GRIN lens 380, another one to acollimator 320, which directs the reference beam 350 to a beam splitter370. The beam splitter 370 directs the reference beam onto the detector360. It is contemplated that other optical components and combinationsthereof, such as lens, prisms, mirrors, fibers, etc., may be used todirect the light from the light source to the illumination path and thereference beam path. Note that although not shown in FIGS. 1 and 2,similar arrangements are contemplated to provide both the illuminationbeam and reference beam.

The backscattered light from the sample is collected by thedetector—either via imaging the scattering volumetric sample or via thedetection of wave fields from scatterers from all depths within thesample. In both cases the reconstruction algorithm developed forholoscopy can be used for data reconstruction.

REFERENCES

1. Mujat, M., Iftimia, N. V., Ferguson, R. D., Hammer, D. X.Swept-source parallel OCT. Proc. SPIE 7168, Optical Coherence Tomographyand Coherence Domain Optical Methods in Biomedicine XIII. February 2009,p. 71681E.

2. Bonin, T., Franke, G. L., Hagen-Eggert, M., Koch, P., Hüttmann, G. Invivo Fourier-domain full-field OCT of the human retina with 1.5 millionA-lines/s. Optics Letters, Vol. 35, Issue 20. 2010, pp. 3432-3434.

3. Hillmann, D., Lührs, Chr., Bonin, T., Koch, P., Hüttmann, G.Holoscopy—holographic optical coherence tomography. Optics Letters, Vol.36, Issue 13. 2011, 2390-2.

4. Hillmann, D., Franke, G., Lührs, C., Koch, P Hüttmann, G. Efficientholoscopy image reconstruction. Optics Express, Vol. 20, Issue 19, pp.21247-21263. 2012.

5. Hillmann, D., Franke, G., Hinkel, L., Bonin, T., Koch, P., Hüttmann,G. Off-axis full-field swept-source optical coherence tomography usingholographic refocusing. Proc. SPIE 8571, Optical Coherence Tomographyand Coherence Domain Optical Methods in Biomedicine XVII. 2013, 857104.

6. Nanjo, T. Fundus camera with partially common coaxial observation andphotographing optical systems. United States Patents, 1996. U.S. Pat.No. 5,543,865.

7. Nunokawa, K. Eye fundus camera having ring slit mask in illuminationsystem. United States Patent, 1981. U.S. Pat. No. 4,422,736.

8. Matsumura, I., Kohayakawa, Y. Eye fundus camera with focus settingdevice. United States Patent, 1974. U.S. Pat. No. 3,925,793.

9. El-Bayadi, G. New method of slit-lamp micro-ophthalmoscopy. Brit. J.Ophthal., Vol. 37, pp 625628. 1953.

10. Kitajima, N., Okamura, K. Slit lamp microscope. United StatesPatents, 2000. U.S. Pat. No. 6,072,623.

While the present invention has been described at some length and withsome particularity with respect to the several described embodiments, itis not intended that it should be limited to any such particulars orembodiments or any particular embodiment, but it is to be construed withreferences to the appended claims so as to provide the broadest possibleinterpretation of such claims in view of the prior art and, therefore,to effectively encompass the intended scope of the invention.Furthermore, the foregoing describes the invention in terms ofembodiments foreseen by the inventor for which an enabling descriptionwas available, notwithstanding that insubstantial modifications of theinvention, not presently foreseen, may nonetheless represent equivalentsthereto.

What is claimed is:
 1. An ophthalmologic imaging system comprising: atunable light source configured to provide an illumination beam and areference beam; a mirror; and an optical system configured to: focus theillumination beam from the light source onto the mirror, such that themirror reflects the light onto a sample in a first light path, and thelight is partly scattered at a desired area of the sample and partlyreflected at a plane or curved interface in the sample; detect scatteredlight from the desired area of the sample by a detector situated in asecond light path, and direct the reference beam onto the detector, thereference beam being off-axis to the second path; wherein an angle ofthe mirror is configured to reflect the light in the first light path,which is at an angle to the second light path; and wherein the mirror isnot in the second light path; wherein the tunable light source has atotal coherence length, defined by the spectral width of the tuningrange, not exceeding 50 μm, and an instantaneous coherence length,defined by the instantaneous line width of the spectral range during thescan, of at least 200 μm; wherein the detector is configured to collectthe reflected light from the sample via imaging the scatteringvolumetric sample or via the detection of wave fields from scatterersfrom all depths within the sample.
 2. The system of claim 1, wherein thedetector has an acquisition speed of at least 500 images per second. 3.An ophthalmologic imaging system comprising: a tunable light sourceconfigured to provide an illumination beam and a reference beam; amirror with a hole; and an optical system configured to: focus theillumination beam from the light source into the hole of the mirror,such that the illumination beam passes through the hole and illuminatesa desired area in a sample through a lens, and the light is partlyscattered at the desired area in the sample and partly reflected at aplane or curved interface in the sample, wherein the scattered lightfrom the desired area in the sample is converted into far field light bya lens in the sample and the reflected light from the plane or curvedinterface in the sample is imaged back through the hole by the lens;direct the far field light onto a detector by the mirror; and direct thereference beam onto the detector, the reference beam being off-axis to alight path from the sample to the detector; wherein the tunable lightsource has a total coherence length, defined by the spectral width ofthe tuning range, not exceeding 50 μm, and an instantaneous coherencelength, defined by the instantaneous line width of the spectral rangeduring the scan, of at least 200 μm; wherein the detector is configuredto collect the reflected light from the sample via imaging thescattering volumetric sample or via the detection of wave fields fromscatterers from all depths within the sample.
 4. The system of claim 3,wherein the detector has an acquisition speed of at least 500 images persecond.
 5. The system of claim 3, wherein the sample is an eye and thedesired area is on the retina of the eye, such that a focus is on orslightly beneath the cornea and the cornea reflex passes through thehole in the mirror, not reaching the detector.
 6. An ophthalmologicimaging method comprising: illuminating a sample by a beam from atunable light source; focusing the beam onto a mirror, such that themirror reflects the light onto the sample in a first light path, and thelight is partly scattered at the sample and partly reflected at a planeor curved interface in the sample; detecting scattered light from thesample by a detector situated in a second light path, and directing areference beam from the tunable light source onto the detector, thereference beam being off-axis to the second path; and configuring anangle of the mirror to reflect the light in the first light path, whichis at an angle to the second light path; and wherein the mirror is notin the second light path; wherein the tunable light source has a totalcoherence length, defined by the spectral width of the tuning range, notexceeding 50 μm, and an instantaneous coherence length, defined by theinstantaneous line width of the spectral range during the scan, of atleast 200 μm; wherein the detector is configured to collect thereflected light from the sample via imaging the scattering volumetricsample or via the detection of wave fields from scatterers from alldepths within the sample.
 7. The method of claim 6, wherein the detectorhas an acquisition speed of at least 500 images per second.
 8. Anophthalmologic imaging method comprising: illuminating a sample by abeam from a tunable light source; focusing the beam onto a hole of amirror, such that the beam passes through the hole and illuminates adesired area in a sample through a lens, and the light is partlyscattered at the desired area in the sample and partly reflected at aplane or curved interface in the sample, wherein the scattered lightfrom the desired area in the sample is converted into far field light bya lens in the sample and the reflected light from the plane or curvedinterface in the sample is imaged back through the hole by the lens;directing the far field light onto a detector by the mirror; anddirecting a reference beam from the tunable light source onto thedetector, the reference beam being off-axis to a light path from thesample to the detector; wherein the tunable light source has a totalcoherence length, defined by the spectral width of the tuning range, notexceeding 50 μm, and an instantaneous coherence length, defined by theinstantaneous line width of the spectral range during the scan, of atleast 200 μm; wherein the detector is configured to collect thereflected light from the sample via imaging the scattering volumetricsample or via the detection of wave fields from scatterers from alldepths within the sample.
 9. The method of claim 8, wherein the sampleis an eye and the desired area is on the retina of the eye, such that afocus is on or slightly beneath the cornea and the cornea reflex passesthrough the hole in the mirror, not reaching the detector.
 10. Themethod of claim 8, wherein the detector has an acquisition speed of atleast 500 images per second.